Magnetic resonance imaging (“MRI”) is a widely accepted and commercially available technique for obtaining digitized visual images representing the internal structures of objects, such as the tissues of the human body, having substantial populations of atomic nuclei that are susceptible to nuclear magnetic resonance (“NMR”) phenomena. In MRI, the nuclei in a structure to be imaged are polarized by imposing a strong, uniform magnetic field, B0, on the nuclei. Selected nuclei are then excited by imposing B1, a radio frequency (“RF”) signal at a predetermined NMR frequency. By doing this repeatedly while applying different magnetic field gradients and suitably analyzing the resulting RF responses from the nuclei, a map or image of the relative NMR responses as a function of nuclei location may be determined. Data representing the NMR responses in space may be displayed.
Referring to FIG. 1, a conventional MRI system 10 typically includes a magnet 12 operable for imposing the strong, uniform magnetic field B0, a plurality of gradient coils 14 operable for imposing the magnetic field gradients in three (3) orthogonal coordinates, and a plurality of RF coils 16 operable for transmitting and receiving RF signals to and from the selected nuclei. The RF coils 16 may be used for transmitting, receiving, or both. The NMR signal received by each RF coil 16 is transmitted to a computer 18 operable for processing the data into an image on a display 20. The MR image is composed of picture elements referred to as “pixels.” The intensity of a given pixel is proportional to the NMR signal intensity of the contents of a corresponding volume element or “voxel” of the structure being imaged. The computer 18 also controls the operation of the gradient coils 14 and the RF coils 16 through a plurality of gradient amplifiers 22 and an RF amplifier/detector 24, respectively.
Each voxel of an image of the human body contains one or more tissues. These tissues contain primarily fat and water, which, in turn, include a plurality of hydrogen atoms. In fact, the human body is approximately 63% hydrogen atoms. Because hydrogen nuclei have a readily discernible NMR signal, MRI of the human body primarily images the NMR signal from the hydrogen nuclei.
In NMR, the strong, uniform magnetic field B0 is employed to align nuclei that have an odd number of protons and/or neutrons, such that the nuclei have a spin angular momentum and a magnetic dipole moment. The magnetic field(s) B1, applied as a single pulse transverse to the strong, uniform magnetic filed B0, pump energy into the nuclei, causing the angular orientation of the nuclei to flip by, for example, 90 degrees or 180 degrees. Following this excitation, the nuclei precess and gradually relax into alignment with the strong, uniform magnetic field B0. As they precess and relax, the nuclei emit energy in the form of weak but detectable free induction decay (“FID”). These FID signals and/or RF or magnetic gradient refocused “echoes” thereof, collectively referred to as MR signals, sensed by the NMR imaging system are analyzed by the computer 18 and used to produce images of, for example, a structure of the human body.
The excitation frequency and the FID frequency are related by the Larmor equation. This equation states that the angular frequency, ω0, of the precession of the nuclei is the product of the strong, uniform magnetic field B0 and the so-called magnetogyric ratio, γ, a fundamental physical constant for each nuclear species:ω0=B0γ.  (1)
By superimposing a linear magnetic field gradient, Bz=Z Gz, on the strong, uniform magnetic field B0, which is typically defined as the Z-axis, for example, nuclei in a selected X-Y plane may be excited by the appropriate choice of the frequency of the transverse excitation field applied along the X or Y axis. Similarly, a magnetic field gradient may be applied in the X-Y plane during the detection of the MR signals to spatially localize emitted MR signals from the selected X-Y plane according to their frequency and/or phase.
Typically, an MRI operator sets a repetition time, TR, and a flip angle of multi-slice gradient echo MR images, or of a given scan. Predetermined TR values allow for the completion of the scan in a minimum amount of time. These TR values depend upon, among other things, the number of slices to be imaged. In general, relatively longer or higher TR values provide a relatively higher signal-to-noise ratio, S/N, in the images. One problem, however, is that changing the TR value may change the image contrast, often in an undesirable way. This contrast may be restored to the contrast associated with an original, desirable scan with a known TR value and flip angle, however this may take relatively more time. In general, it is desirable to minimize MRI set-up and acquisition times, while maintaining a relatively high S/N and a contrast similar to that of a target image. Thus, the conventional trial-and-error approach to selecting a TR value and a flip angle is undesirable.
Elaborating on the above, it is often desirable to change the TR of a sequence of MR images, preferably without changing the contrast of the images. Typically, this is done to reduce the time required to complete a scan. For example, given a 256-matrix with a TR of about 0.125 seconds, an imaging time of about 32 seconds is required. Further, given the fact that an echo time, TE, and other necessary delays are such that the MRI system has an inter-slice delay of about 0.025 seconds, as many as five (5) slices may be imaged simultaneously, but no more.
Suppose that the simultaneous imaging of six (6) slices is desired. Because six slices may not be fit into the TR of about 0.125 seconds, the additional slice requires repeating the scan, reducing the time efficiency of the scan to approximately 60%. However, increasing TR by a factor of about 6/5 allows all of the slices to be imaged in about 38 seconds. In other words, increasing TR to about 0.15 seconds allows the desired images to be produced approximately 67% faster than is possible at a TR of about 0.125 seconds.
Suppose that the imaging of fifteen (15) slices is desired. Assuming the same TR as above, it takes about 96 seconds to image the fifteen slices. Five (5) slices are imaged during the first 32 seconds, five (5) slices are imaged during the second 32 seconds, and five (5) slices are imaged during the third 32 seconds. Alternatively, the fifteen slices could be imaged in the same 96 seconds, fifteen at a time, if the TR is tripled. Tripling the TR, in this case, would provide a considerable increase in S/N.
The relationship between signal intensity and relaxation determines the image contrast (T1 contrast or T1-weighting in the case of longitudinal relaxation). Although there is no universally accepted definition of T1-weighting, an image is said to be strongly or heavily T1-weighted if the T1 differences among imaged regions lead to large intensity differences among the regions. It is often desirable, for example, to minimize this T1-weighting.
Because the signal intensity associated with an image is proportional to the longitudinal magnetization present immediately prior to the introduction of the RF excitation pulses, Mz, a relatively simple equation for Mz is typically used to predict contrast:Mz=(1−exp(−rTR))/(1−cos(flip)exp(−rTR)).  (2)(See also Equation 5). Here, Mz is normalized to “1” following an infinite wait and the relaxation rate, r, is 1/T1. The effects of T2 and T2* have been stripped from this equation. The flip angle is assumed to be constant across the thickness of the excited slice. Often, the TR may be varied without significantly affecting the contrast. Increasing the TR at a constant flip angle changes the Mz versus r plot significantly, as does increasing the flip angle at a constant TR. However, a coordinated change in the TR and the flip angle has a relatively small effect on the contrast. Thus, a method is needed to determine the appropriate flip angle to produce a desired contrast at a desired TR, and to choose the appropriate TR should a different flip angle be desired.
Increasing the TR may increase S/N, particularly when the flip angle is also increased. Consider again the fifteen (15)-slice case. Tripling the TR while increasing the flip angle from 20 to 35 degrees leaves Mz and the contrast essentially unchanged. The largest Mz difference these changes may cause, for example, is 0.035 at T1=⅓ seconds. Starting with the same Mz, the 35 degree pulse produces relatively more signal by the ratio sin(35 degrees)/sin(20 degrees), or 1.677 (where Mz is not substantially preserved, as for relatively large TR changes, the ratio of sines does not accurately indicate signal change).
Thus, what is needed are automated systems and methods that allow an MRI operator to quickly and easily set up and complete a scan using a computer, incorporating parameters, including TR values and flip angles, from a previous scan that has a desired contrast. This would allow the MRI operator to obtain high-S/N images with the desired contrast in the minimum amount of set-up and scan time.